Not applicable.
The present invention relates to imaging systems and more particularly to an imaging system including at least two detectors, a first detector being a high resolution detector which detects energy and location of Compton scatters and a second detector being a residual energy detector.
Radionuclides are decaying substances which emit subatomic particles (e.g. beta particles, alpha particles, neutrons, positrons and/or photons). For example single photon emitters such as Technetium emit photons and perhaps also a charged particle wherein there is no angular correlation between multiple emitted photons. As another example, Fluorine-18 decays to Oxygen-17 by emitting a positron which has an energy between 0 and 600 keV with a mean energy of 250 keV.
Radionuclides are employed as radioactive tracers called xe2x80x9cradiopharmaceuticalsxe2x80x9d by incorporating them into substances such as glucose or carbon dioxide. One common use for radiopharmaceuticals is in the medical imaging field. To use a radiopharmaceutical in imaging, the radiopharmaceutical is injected into a patient and accumulates in an organ, vessel or the like, which is to be imaged. Hereinafter an exemplary radionuclide in an exemplary radiopharmaceutical used for imaging will be referred to as an imaging radionuclide.
It is known that specific radiopharmaceuticals become concentrated within certain organs. The process of concentrating often involves processes such as glucose metabolism, fatty acid metabolism and protein synthesis. Hereinafter, an organ to be imaged will be referred to generally as an xe2x80x9corgan of interestxe2x80x9d and prior art and the invention will be described with respect to a hypothetical organ of interest.
After a radiopharmaceutical becomes concentrated within an organ of interest and while the imaging radionuclides decay, the radionuclides emit subatomic particles including positrons and photons. Each of the particles and photons can be detected. Where the particles are positrons, the positrons travel very short distances (e.g., approximately 200 microns in the case of Flourine-18) before they encounter an electron and, when a positron encounters an electron, the positron is annihilated as the electron and positron unite and two photons are generated.
Photons are characterized by one feature which is pertinent to all medical imaging techniques which sense photons. Given a specific radionuclide, photons generated directly from decay or via decay followed by annihilation, have specific and known energy levels. For example, when a positron results from decay of Fluorine-18, annihilation of the positron always results in two photons, each of which has an energy of 511 keV. As other examples, 131I generates photons having energies of 360 keV while 99mTc generates photons having energies of 140 keV. In addition, particle annihilation events are characterized by an additional feature which is pertinent to medical imaging. This additional feature is that, upon an annihilation event, two photons are generated and the photons are directed in essentially opposite directions (i.e. the trajectories are separated by approximately 180xc2x0). There is approximately a xc2x10.25 degree variation from 180xc2x0 in the photon trajectories related to the momentum of the electron-positron pair before annihilation.
In all photon imaging systems both photon energy and trajectory have to be determined. Photon energy is determined and compared to a range of expected energies associated with the particular radiopharmaceutical used during data generation. Where sensed energy is outside the expected range the detection is typically discarded. Where sensed energy is within the expected range, the detection is identified as valid, photon trajectory is determined and trajectories for all valid detections are combined to generate an image of the object of interest.
Three different imaging systems which are pertinent to the present invention include PET, collimated single photon imaging and Compton imaging systems, each of which is described separately below.
PET Systems
An exemplary PET system includes two oppositely facing cameras wherein the cameras are either scintillation cameras or solid state direct conversion detector (DCD) cameras.
An exemplary scintillation camera includes a plurality of detector units and a processor which, among other things, includes coincidence detection circuitry. An exemplary detector unit includes a two dimensional 6xc3x976 matrix of bismuth germinate (BGO) scintillator crystals which are disposed in front of four photo multiplier tubes (PMTs). When a crystal absorbs a photon, the crystal generates light which is generally directed toward the PMTs. The PMTs absorb the light and each PMT produces an analog signal which arises sharply when a scintillation event occurs and then tails off exponentially with a time constant of approximately 300 nanoseconds. The relative magnitudes of the analog PMT signals are determined by the position in the 6xc3x976 BGO matrix of the crystal which generates the light (i.e., where the scintillation event takes place), and the total magnitude of these signals is determined by the energy of the photon which causes an event.
For each total magnitude within a range of expected magnitudes corresponding to the imaging radionuclide, a set of acquisition circuits receives the PMT signals and determines x and y event coordinates within the BGO matrix thereby determining the crystal which absorbed the photon and the general x-y coordinate at which the absorption occurred on the face of the crystal. Each acquisition circuit also produces an event detection pulse (EDP) which indicates the exact moment at which a scintillation event took place.
The information regarding each valid event is assembled into a digital event data packet which indicates precisely when the event took place and the position of the BGO crystal which detected the event. Event data packets are conveyed to a coincidence detector which determines if any two events from the opposing detectors are in coincidence.
Coincidence is determined by a number of factors. First, the time markers in each event data packet must be within a specific time window of each other, and second, the locations indicated by the two event data packets must lie on a straight line which passes through the field of view of a scanner imaging area. Events which cannot be paired as coincidence events are discarded, but coincidence event pairs are located and recorded as coincidence data packets. Each coincidence data packet includes a pair of digital numbers which precisely identify the addresses of the two BGO crystals that detected the event. After an event pair has been identified, the source location of the pair can be identified along a straight line which passes through the locations of the events in the pair. After imaging data has been collected in this manner, a processor uses the collected information to generate a two or three dimensional image of the organ of interest.
DCDs may be based on pixilated semiconductor detectors such as Cadmium Telluride (CdTe) or Cadmium Zinc Telluride (CdZnTe) devices. Generally, each DCD includes an absorption member, a cathode, at least one anode, a potential biasing mechanism (i.e. voltage source) and a separate amplifier for each anode.
The absorption member is formed of a planar semiconductor material (e.g. CdTe or CdZnTe) which has oppositely facing cathode and anode surfaces. The dimension between the cathode and anode surfaces is an absorption member thickness. When photons are directed at the cathode surface, the photons penetrate the absorption member and each photon is absorbed at an absorption depth within the member thickness. When a photon interacts with the absorption member while being absorbed, the absorption member generates a plurality of electrons and holes.
The cathode is attached to and essentially covers the cathode surface and the anode is attached to the anode surface. The biasing mechanism is linked to the cathode and biases the cathode negative. The anode remains unbiased and therefore is positive with respect to the cathode. Because the cathode is negative and the anode is positive with respect to the cathode, when electrons and holes are generated during absorption, the holes are attracted to the cathode surface and the electrons are attracted to the anode surface. The electrons generate a first negative charge component on the anode.
As holes accumulate at the cathode, the positive charge adjacent the cathode causes a capacitive second negative charge component on the anode. To distinguish between the first negative charge component on the anode caused by electrons which travel from the absorption depth to the anode and the second negative charge component on the anode caused by the holes, the first negative charge component will be referred to hereinafter as the electron charge and the second negative charge component will be referred to hereinafter as the hole charge. Together, the electron charge and the hole charge are referred to hereinafter as the collected charge.
The amplifier is attached to the anode and includes an output lead for providing an anode signal indicating the collected charge. The amplifier output lead is linked to a camera processor. The processor integrates the anode signal over an integration period and provides an intensity signal. The processor compares the intensity signal to an expected intensity signal or expected energy range (e.g., 511 keV) associated with the imaging radionuclide. When an intensity signal is within the expected range, the processor indicates that a photon has been detected by the DCD which provided the anode signal.
As with the scintillation camera, DCD cameras provide photon detection signals to coincidence circuitry which in turn identifies coincident pairs of photons and stores the coincident pairs as coincident data packets for subsequent image processing.
Two important criteria for any imaging system are resolution and sensitivity. Resolution is a term used to refer to position accuracy of a sensed interaction or energy deposit. In other words, resolution measures how close a perceived absorption point is to an actual absorption point. Sensitivity is a term used to refer to the percentage of photons within an expected energy range emanating toward a camera which are actually detected to be valid events. High sensitivity is better than low sensitivity. For example, a sensitivity of 20% (i.e., 1 in 5 photons are detected) is better than a sensitivity of 10%.
The resolution criteria favors selecting a radiopharmaceutical which generates relatively low energy photons. Three sources of resolution degradation include Compton scattering, depth-of interaction variances and incident angle errors. As well known in the imaging arts, in addition to complete absorption, a second type of interaction referred to as xe2x80x9cCompton scatteringxe2x80x9d often takes place within an absorption member or scintillation crystal. Hereinafter, the term xe2x80x9cabsorption memberxe2x80x9d will be used generically to refer to either a solid state absorption member or a scintillation crystal. When a photon enters an absorption member, the photon may experience a first interaction in which photon direction is altered and only a portion of photon energy is absorbed. Thereafter, the photon may exit the absorption member without being fully absorbed, Compton scatter one or more additional times prior to full absorption, or may be fully absorbed at some other location within the member upon a second interaction.
To increase the likelihood of full absorption and hence the sensitivity of a detector, most PET detectors are designed to have relatively thick absorption members. In this manner, while a photon may xe2x80x9crattlexe2x80x9d around in the member from one Compton scattering to the next prior to complete absorption, the entire photon energy will be sensed and therefore the energy will be within the expected energy window.
Where photon energy is absorbed at several different locations (i.e. one or more Compton scatters occur prior to final absorption), the location of the first interaction or absorption is difficult and, in some cases, impossible to determine. For instance, referring to FIG. 1, an exemplary absorption member 10 is illustrated which includes an entry face 11. A photon 12 emanates from an object of interest (not illustrated) and travels along a path 14 into member 10 through an entry point X in face 11 and a single Compton scatter occurs at point A. The scattered photon traverses along a path essentially parallel to face 11 prior to a complete absorption of residual photon energy at point B. In the case of these two interactions the entry point is perceived to be located at the xe2x80x9ccenter of gravityxe2x80x9d of the two energy depositions. For example, if deposition B is much greater than deposition A, the perceived entry point is close to Y. The potential for error increases as the number of Compton scatters increases.
With respect to depth-of-interaction errors, referring still to FIG. 1, an absorption may occur at any depth within the thickness of absorption member 10. The coordinate detection circuitry is set up to identify an x-y coordinate pair (i.e. a point on face 11) corresponding to the center of gravity of all related depositions. This type of detection works well for photons which enter absorption member 10 perpendicular to face 11 but results in position errors where photons enter the member at an angle with respect to face 11. For example, assume a photon 16 enters member 10 at a 45 degree angle with respect to face 11 at point C and traverses along a path 17 to point D prior to full absorption. In this case instead of identifying point C as the location at which photon 16 entered face 11, detector circuitry identifies a face entry point E associated with point D. Once again an error occurs.
With respect to incident angle errors, in PET systems, as indicated above, while photons generated by a single annihilation travel essentially in opposite directions (i.e. along trajectories which are 180xc2x0 apart), there is some variation (i.e. xc2x10.25 degrees) from the 180xc2x0 assumption. Because of this variation, the location of an annihilation identified via a PET system often has some slight error. For example, where PET cameras are one meter apart, 0.25 degrees variation translated into xc2x12 mm halfway between the detectors.
As well known in the imaging industry, when a radionuclide which generates low energy photons is used for imaging, the probability of interaction within an absorption member is relatively high, absorption typically occurs upon a first interaction and the first interaction typically occurs at a relatively shallow depth within the absorption member. For these reasons Compton scatter and depth-of-interaction variances do not appreciably effect quality when low energy photons are employed.
As incident photon energy increases, the probability of a complete absorption drops very fast while the probability of Compton scattering drops more slowly. Consequently, at 511 keV, for example, a large fraction of photons which are completely absorbed within a member will Compton scatter during a first interaction and are subsequently absorbed somewhere else within the member. In addition, at 511 keV, photons are likely to travel tens of millimeters in most detectors prior to interaction and therefore depth-of-interaction variances tend to distort final images.
In any event, Compton scattering, depth of interaction variances and incident angle variances combined when imaging with medium to high energy photons result in relatively poor (e.g. 3-5 min) imaging resolution. While such resolution may be sufficient where images of large objects are to be generated, much higher resolution (e.g. 500 microns or below) is required when small animals (e.g. a mouse) or a limited region of interest in humans is to be examined.
Although imaging of radiotracers labeled with positron emitting radionuclides is extremely useful, positron emitters tend to have short half lives (e.g., C-11 (20 seconds), 0-15 (2 minutes), F-18 (2 hours)). This makes such radiotracers unsuitable for studies in which the specific radiopharmaceutical is known to take several days to concentrate in the organ of interest. Furthermore, some compounds of biological interest may not be readily labeled with available positron-emitting nuclides, or the resulting radiopharmaceutical may possess undesirable biochemical characteristics. It is also the case that short lived positron emitters must be produced on site, and this often requires an expensive cyclotron installation and radiopharmaceutical preparation facility.
PET imaging also has other inherent shortcomings. First, PET imaging depends upon detection of both photons from a single annihilation event and therefore depends upon the entire path length through any intermediate attenuator (e.g., the mass of a patient between an annihilation location and the detector). In fact, it has been observed that attenuation in PET coincidence mode at 511 keV for a source at the center of an object is always greater than the attenuation of even 90 keV photons in a single photon counting mode. This severe attenuation results in few photons detected from xe2x80x9cdeepxe2x80x9d structures in PET.
Second, in volume PET imaging, sensitivity decreases significantly toward the edges of the axial field-of-view with the joint angle xe2x80x9cseenxe2x80x9d by the annihilation radiation. In other words, often one annihilation photon in a pair may be detected while the other photon in the pair shoots axially out of an imaging area and is never detected.
Third, even where an annihilation pair reaches opposing PET detector sections, often only one of the two photons will be detected, the other of the two photons passing through the absorption member without an absorption event.
Fourth, where more than two absorption events are simultaneously detected with a conventional PET system there is no good way to determine which of the multiple events are associated with a single annihilation and hence there is no way to determine the sources of the photons. In effect, the data is lost.
Collimated Single Photon Camera
There are many radionuclides which emit one or more photons which are not correlated in angle as are the annihilation photons related to positron emitters. These are known as single photon emitters. Single photon emitters are readily available with a wide range of chemical properties, photon energies and half-lives. Single photon emitters cannot be imaged using coincidence techniques in the same manner as positron emitters.
Another imaging system which relies on detection of emissions and can detect single photons is a mechanically collimated emission camera. A collimated camera is similar to the construction of a single PET camera in that this type of camera includes some type of absorption member which is capable of sensing a photon""s absorption energy and location. To determine the angle of photon flight prior to absorption a collimated camera includes a collimator which essentially restricts absorbed photons to known paths which are often perpendicular to a broad face of the absorption member. While a collimated camera reduces the amount of calculations required to identify the source of a detected photon, collimated cameras have extremely low sensitivity and resolution is negatively affected by the collimator.
Compton Camera Imaging
An exemplary Compton camera includes first and second detectors which are both arranged to one side of an imaging area and the position of the second detector with respect to the first is locked and known. The first detector is designed to cause a photon to Compton scatter (i.e., a scattering event) within a scattering member so that photons emanate from the first detector along modified trajectories and having modified energies. The first detector senses the position of the scattering event, the energy absorbed during the scattering event and the time of the scattering event. To this end the first detector is typically relatively thin so that the number of photons which are completely absorbed within the first detector is relatively small. Related effects of a thin detector include reduced interactions after a first interaction and a smaller range of depth-of-interaction variances.
The second detector is configured and positioned such that the detector is within a path likely to be traversed by scattered photons. To this end, as scattered photons may scatter in virtually any direction, most second Compton camera detectors define a space in which the first detector is positioned. For instance, an exemplary second detector may have the shape of a box with an open face, the first detector being positioned within the open face so that any photon which scatters from the first detector into the box, despite the angle of scatter, will be detected by the second detector.
An absorption member within the second detector absorbs the scattered photons (i.e., an absorption event), identifies the positions of the absorption events, the energies absorbed during the absorption events and the times of the absorption events. The energies and locations of coincident scattering and absorption events are combined with a knowledge of the expected energy of photons generated by the imaging radionuclide to identify, within a conical ambiguity, the possible paths of a corresponding photon prior to collision with the first detector.
After conical data corresponding to a large number of detected photons has been generated, tomographic techniques are employed to locate the origin of the photons by finding the intersections of many different possible path cones corresponding to different detected photons. As in the case of PET, the source data is then combined to generate an image of the object of interest.
The quality of images generated using a Compton camera, like the quality using a PET system, is degraded by both multiple Compton scattering and depth-of-interaction variances in the second detector. In fact, Compton camera systems which depend upon scintillation detectors for the second detector have even a worse problem with uncertainty in the first detector interaction point than conventional PET cameras. In addition, such Compton cameras also generally have lower sensitivity than PET cameras.
Moreover, Compton cameras also have a number of additional shortcomings. Specifically, Compton cameras require extremely accurate first and second detectors, each of which can provide accurate event times, energies and locations. In the case of the first detector, required accuracy is not particularly burdensome as the area of the first detector is relatively small and hence the cost of configuring an accurate detector is practical. However, in the case of the second detector, the area of the second detector is relatively large (i.e., the entire internal surface of a box shaped detector) and hence the associated costs are appreciable.
Second, the processor required to resolve the conical ambiguities among many different valid events has to be extremely computationally capable. Such processors are relatively extensive when compared to the processors required to manipulate PET data.
In addition to the problems discussed above, radionuclide imaging generally has a number of other shortcomings. First, as well known in the imaging art, image quality can be increased by reducing the distance between a photon source and a detector. For instance, in the case of a Compton camera, the closer the first detector is to the object of interest, the smaller the spatial uncertainty corresponding to a given angular uncertainty. Moreover, being close to the object of interest increases the solid angle subtended by the first detector from the point of interest and camera sensitivity increases correspondingly. For this reason Compton cameras are typically mounted on multi-articulate arms which can position the first detector adjacent the object of interest. Unfortunately, because the arm must support each of the first and second detectors, the arm must be relatively large and often complicated. In addition, even with a suitable arm, often the external surface of the object of interest or the body in which the object resides is much different than the surface of the first detector so that distances between the object of interest and the first detector are appreciable.
Second, often, in addition to being absorbed by an object of interest, a radiopharmaceutical will be absorbed by other tissues or organs which are not of interest but which are proximate the organ of interest. In this case it may be difficult to differentiate between photons emanating from the organ of interest and photons emanating from surrounding tissue or organs. One solution may be to block photons from surrounding tissue and organs using radio-opaque shielding (e.g., a lead shield) or a collimator. Unfortunately, in cases where there is some distance between the organ of interest and the first detector, a blocking shield is relatively ineffective as photons from the tissue and organs adjacent the organ of interest can impact the first detector on an angle. In addition, where the organ of interest is not proximate the first detector, an effective collimator which could block angled photons from the tissue and the organ of interest would have to have extremely small apertures. Such a collimator would substantially reduce the sensitivity of a Compton camera thereby minimizing one of the advantages typically associated with Compton systems.
Because there are several shortcomings and advantages associated with each of the systems described above, it is advantageous for any medical or other type of facility which employs emission imaging systems to have one of each imaging system. Unfortunately, each imaging system is extremely expensive and therefore most facilities have been forced to chose one system and its advantages and shortcomings over the other systems.
One solution which has enabled both PET and single emission imaging using a single system includes two collimated single emission cameras including a coincidence processor. Each of the cameras can be used separately for single emission imaging. In the alternative, the cameras can be arranged so as to oppose each other, can be linked to the coincidence processor and the collimators can be removed so that a PET system is configured. No similar xe2x80x9cdouble dutyxe2x80x9d system has been provided for PET and Compton imaging.
A need exists for a PET imaging system which is extremely accurate when a radiopharmaceutical which generates high energy photons is used to generate imaging data. In addition, it is always advantageous to have a Compton camera which includes a first detector which is as close as possible to an object of interest and therefore any Compton camera configuration which can reduce the distance between an object of interest and a first detector would be a welcome development. Moreover, an imaging system which could facilitate both Compton and PET imaging and which could increase imaging sensitivity at minimal expense would be particularly advantageous.
In emission imaging systems photons which emanate from a radionuclide concentrated within an object of interest are detected and used to generate an image of the object. To this end two important detected photon characteristics must be determined for the detected photon to be useful for imaging purposes. First, the photon energy must be determined and compared to the expected energy of photons generated by the imaging radionuclide. Where a photon is not within the expected range either the photon is from a source other than the radionuclide in the object or the photon""s energy was reduced by some interaction (e.g., collision with other matter in the patient""s body or within an imaging vicinity) which likely changed the photon trajectory prior to detection. In either of these cases, the photon should not be used for imaging and should be discarded. Second, for photons within the expected energy range, the photon path must be determined so that the path can be traced back to the photon source.
As indicated above, upon a first interaction within an absorption member high energy photons deposit some energy and then tend to Compton scatter in a random direction. Nevertheless, if a single scattering event occurs at an event location and then the scattered photon exits the detector, the event location is accurate.
By providing opposing detectors on opposite sides of an imaging area, each of which causes a scattering event when two photons from a single annihilation are detected, as with any PET imaging system, the location of the annihilation and hence the location of the photon source can be determined as being along a line between the two scattering event locations.
Unfortunately, because the scattering detector scatters photons, there is no way, with the scattering detectors alone, to determine the total energy of the sensed photon. For this reason, with the scattering detectors, there is no way to determine if the photon is from the imaging radionuclide or whether the photon changed trajectory after the annihilation.
According to the present invention a second detector can be provided for each first detector wherein the second detector is designed to totally absorb scattered photons thereby determining the residual energy of each scattered photon. Thereafter, the scattering event energy and the residual or absorption energy can be combined to determine the total or sensed energy corresponding to the detected photon. Then, the sensed energy can be used to determine if the sensed photon corresponds to a valid event (i.e., has an energy within the expected energy range given the imaging radionuclide).
Thus, by providing opposing cameras wherein each camera includes a scattering detector and an absorbing detector, scattering and absorption energies can be combined to identify valid sensed events and then coincident scattering event locations for valid events in the opposing detectors can be identified for determining photon source location and for additional imaging purposes.
Thus, one object of the invention is to provide a relatively inexpensive imaging system. To this end, the second detector in each camera need not identify absorption event location. In addition, conventional PET processing as opposed to Compton type processing can be used to determine photon source location.
Another object is to provide an imaging system which can be used with high energy photons. The inventive system works best with high energy photons which are more likely to Compton scatter than to be absorbed upon an interaction with the first detectors.
One other object is to reduce the adverse effects of depth-of-interaction variances. To this end, the first detector in each camera is designed such that the detector can provide an accurate three dimensional scattering location for each scattering event.
Consistent with the objects of the invention, an exemplary embodiment of the invention includes an apparatus for identifying the location of a photon source within an imaging area and which generates photons having energies within a known energy range. The apparatus including two oppositely facing detector pairs or cameras disposed on opposite sides of the imaging area, each camera including a first detector unit which causes Compton scattering when a photon enters the unit and generates signals indicative of the scattering event location, energy and time and a second detector unit which absorbs the scattered photon and generates signals indicative of the absorption event energy and time. Where combined coincident scattering and absorption event energies detected by a camera are within the expected energy range, the scattering event location is identified as the location of a valid event. Thereafter, the times of valid events in each of the opposing cameras are compared and the event locations corresponding to coincident valid events are identified and stored as coincident event pairs. The coincident event pairs can be used to identify the location of the photon source and also can be used subsequently for imaging purposes. The invention also includes a method to be used with the inventive apparatus.
Yet another object of the invention is to provide a Compton camera wherein the first detector can be as close as possible to an object of interest which is to be imaged. To this end, it has been recognized that the first detector in a Compton camera can be very compact and light weight as a Compton camera does not require a bulky mechanical collimator. In addition, it has been recognized that there is substantial leeway in the location of the second detector in the Compton camera. These two factors make it possible to design special purpose imaging geometries where the first detector can be placed very close to the regional anatomy (e.g., breast, prostate, extremity) being imaged, and can be shaped to conform to a patient""s anatomy. Specifically, the first detector can take any desired shape and be placed within a conventional xe2x80x9cimaging areaxe2x80x9d while the second detector(s) can be positioned outside the imaging area.
One other object is to provide a system where the relative positions of the first and second detectors in a Compton camera can be modified while still enabling collection of data meaningful for imaging purposes. To this end, in one embodiment the invention includes a relative position determiner which determines the relative positions of the first and second detectors and the relative position information is then used, along with conventional Compton camera data, to determine photon sources and generate suitable images.
One other object of the invention is to provide a xe2x80x9cCompton probexe2x80x9d which can be used to locate radionuclides and hence tissue and organs which absorb radionuclides within a patient. After removal of a tumor, often residual tumorous tissue may remain. Compton imaging can be used for locating the residual tissue but going in and actually identifying the tissue for retrieval purposes is often a difficult task. According to the invention a small Compton camera is positioned at a distal end of a probe member, the camera including an entrance window which leads into a first detector and then to a second detector. The camera is linked to a processor which receives first and second detector data and can use the data to generate an image of the area in front of the window. To correlate the imaging data as the probe member is moved about, the probe member includes an orientation tracking device which determines the position and orientation of the camera in real time. The processor is linked to a display for displaying an image of the area adjacent the probe window. In effect, a probe user can xe2x80x9cpaintxe2x80x9d a picture of the area adjacent the window by moving the distal end of the probe about proximate a photon source.
One other object is to identify proximity of the inventive probe member to tumorous tissue. As indicated above, positrons only travel a very short distance prior to annihilation. Thus, by providing a positron sensor within the window of the Compton camera at the end of the probe member, when positrons are sensed, tumorous tissue (i.e., tissue with a radionuclide absorbed therein) is located.
Yet another object is to provide a collimator on a first Compton detector which can block photons from tissue and organs which are not of interest but which only minimally reduces camera sensitivity with respect of an object of interest. To this end, it has been recognized that by providing first detectors which are anatomically shaped and therefore can be positioned extremely close to an organ of interest, the fields of view for viewing an organ of interest can be reduced appreciably using shielding or a collimator with minimal sensitivity reduction to photons emanating from the organ of interest.
One other object of the invention is to provide a relatively inexpensive xe2x80x9cmulti-purposexe2x80x9d imaging system which can be used for Compton imaging and PET imaging or can be used in a dual mode wherein both Compton and PET features are combined to increase both sensitivity and source resolution. To this end, it has been recognized that the second detector in a conventional Compton camera is structurally and functionally very similar to detectors used for PET imaging. Thus, in addition to being used for PET imaging, a PET camera can also be used in conjunction with a scattering Compton detector to facilitate Compton imaging wherein the additional hardware expense for Compton imaging is only the expense of the scattering detector.
Furthermore, it has been recognized that in addition to facilitating both Compton and PET imaging/processing, a system including scattering detectors and PET type second detectors arranged in an opposing PET formation can be used to increase sensitivity and resolution. To this end, it has been recognized that a PET system including scattering first detectors can result in three different types of xe2x80x9cusefulxe2x80x9d detected events. The three types of events include conventional PET coincidences where a pair of photons corresponding to a single annihilation event are detected in the PET detector, a single Compton event where one photon corresponding to an annihilation pair escapes detection and the other photon in the pair interacts by Compton scattering in the first detector and absorption in time-coincidence in the second detector and mixed events where the two annihilation photons (which may correspond to either one or two annihilations) have been detected and additionally, one or both photons have scattered in the first detector and are detected in the second detector.
With respect to PET coincidences, these events are processed according to conventional PET techniques to identify source along a line segment connecting the locations of the two interactions. With respect to single photon Compton events, these events are processed according to conventional Compton techniques. With respect to mixed events, these events are processed in any of several different ways depending upon the nature of the detected events. For example, if two Compton scatters corresponding to a single annihilation occur and each scattered photon is absorbed in the second detector, the Compton position information is used in a PET fashion to identify source location. As another example, Compton data can be used to indicate whether a recorded event is a random coincidence or not. For instance, assume each photon in an annihilation pair Compton scatters but only a single scattered photon is absorbed in the second detector. In this case, coincidence can be used to determine that it is likely that the scattered photons form a pair but the energy of the unabsorbed photon cannot be determined because that photon was not absorbed. Compton processing can be applied to scattering and absorption data corresponding to the absorbed photon to identify the source of the absorbed photon within a conical ambiguity. Once the cone of possible photon trajectories is known, if the line between the scattering events is on the cone, it can be assumed the scattering photon which was not absorbed is not a random coincidence and therefore the photon source, along a line, can be determined. Where two coincident events are not from a single annihilation, the events can be Compton processed separately.
A similar procedure can be used to determine if an annihilation photon scattered prior to detection. For example, assume annihilation photons both scatter and are absorbed. In this case a line between scattering event locations identifies a likely source location. To verify source location, Compton processing can be performed on scattering and absorption data corresponding to a single one of the photons to identify source within a conical ambiguity. If the line between scattering events is not on the identified cone, the events are random coincidences.
These and other objects, advantages and aspects of the invention will become apparent from the following description. In the description, reference is made to the accompanying drawings which form a part hereof, and in which there is shown a preferred embodiment of the invention. Such embodiment does not necessarily represent the full scope of the invention and reference is made therefor, to the claims herein for interpreting the scope of the invention.